Computed tomography techniques were first suggested in the 1960s, with practical implementation beginning in the 1970s. The essential principle is that a number of projections are obtained from a number of rotational directions around a single axis of rotation, showing the x-ray attenuation after passing through the object under investigation. Computational techniques are applied to this plurality of projections, to yield a three-dimensional image of the interior of the object. Contrast in the image is derived from the different attenuation rates of the different materials making up the object, and the overall image quality is dependent on the provision of an adequate number of projections. The basic process is set out in U.S. Pat. No. 3,106,640 but has been developed considerably since then.
Typically, a CT scanner will comprise an x-ray source mounted in a rotateable manner around an axis, such as on a ring or a gantry, together with either a single detector mounted opposite the source or a plurality of detectors arranged around the ring. The scanner will be rotated around the axis and will emit pulses of radiation at a predetermined frequency, i.e. with a predetermined time period between them. These pulses will then be detected after attenuation and the resulting series of projections used to compute an image.
The source may be a fan beam directed toward a linear array of detectors, or a cone beam directed towards a two-dimensional detector array. Often, a dedicated investigative CT scanner will use a fan beam illuminating a linear array in order to yield a high-quality image. Such scanners often rotate at a high speed around the patient (or object) under investigation in order to produce an image within a short period of time and to minimise movement artefacts in the image.
Other CT arrangements include a cone-beam arrangement mounted on or as part of the gantry of a radiotherapy apparatus, with the aim of combining radiotherapeutic treatment with obtaining a CT scan. The results of the CT scan can then confirm accurate positioning of the patient and/or guide the radiotherapy delivery. In such cases, the rotational speed of the CT scanner is often dictated by the rotational speed of the radiotherapy gantry, and may be as low as 1 rpm. Such combined systems may use a separate kV x-ray source of CT mounted (for example) 90° away from the therapeutic source, or may rely on a single source able to switch between diagnostic kV emissions and therapeutic MV emissions. Sometimes, a limited form of CT (“portal CT”) is possible using images derived from the therapeutic beam after attenuation by the patient, but the overall contrast of such images is poor given that there is less difference in attenuation coefficients between different materials at the very high energies involved in the therapeutic beam.
Our earlier patent application WO2012/103901 described such a CT system which dealt with the problem that activation of the diagnostic source at a steady frequency might result in unnecessary dosage being delivered to the patient, given that rotation of the gantry was dependent on the needs of the therapeutic source. At times, the rotation rate might be very low or zero, resulting in the production of projection images that were by and large redundant. Our application therefore explained that the diagnostic source should be gated to wait for a minimum angle of rotation to take place between successive images. Gantries are provided with built-in electronics to detect the gantry angle, so the information is readily available.
CT scanning offers good resolution between markedly different material types, such as between bone and soft tissue. However, the attenuation of x-rays by different types of soft-tissue is very similar and therefore CT scans tend to offer poor contrast within soft-tissue areas. One way of addressing this is to employ dual-energy scanning, in which a first set of projection images are obtained at a first x-ray energy such as 80 kVp and a second set are obtained at a second and different x-ray energy such as 140 kVp. These two energies are chosen as they are often the upper and lower limits of the energies available for selection. The x-ray attenuation of materials varies with the x-ray energy, as noted above in relation to portal CT, and this difference in energy is enough to allow different materials to be distinguished by their different differential attenuation. The resulting volumetric image can therefore be marked to identify areas of different materials, such as by artificial colours or by tagging them so that specific material types can be highlighted or removed from the image. A summary of the development of dual-energy CT scanning techniques can be found in the article “An Introduction to Dual Energy Computed Tomography” by Michael Riedel, University of Texas Health Science Center at San Antonio published online at http://ric.uthscsa.edu/personalpages/lancaster/DI2_Projects_2010/dual-energy_CT.pdf.
Dual energy CT scanners therefore typically comprise two kV tubes, operating at two different energies, and mounted on the same rotating mount by separated by a suitable angular separation. Rotating at high speed around the patient, a set of projection images are obtained at both energies. Older dual energy systems employed a single source that made multiple passes around the patient at different energies, but suffered from difficulties arising from natural movement of the patient between passes. The successful implementation of dual energy CT scanning therefore had to wait for the ability to integrate multiple sources into a single apparatus.